1. Technical Field
The present technology pertains generally to systems and methods for the detection and imaging of ionizing radiation, and more particularly to a multiple spatial resolution modular detector block with scintillator compositions and related structures for measuring depth of interaction activity, crystal scatter identification, and improved detection of scintillation events from both high energy and low energy radioisotope distribution fields. The detector blocks can be adapted for use with a variety of detectors and are particularly suited for high performance positron emission tomography.
2. Background
Single crystal scintillation is a very simple and sensitive method for detecting high energy radiation such as x-rays, gamma-rays and high energy particles with energies that exceed a few kilo-electron volts (KeV). Crystals with high light yield, narrow energy resolution and fast decay times are required for medical imaging systems such as positron emission tomographs (PET).
Radiation interacts with a scintillation crystal transforming the energy of the absorbed quanta into multiple photons of scintillation light. The amount of light that is emitted is proportional to the energy of the charged particle and the quantity of ionizing radiation that impinges on the scintillating material.
In the case of gamma radiation, the amount and spatial distribution of light that is produced from the scintillator is dependent on whether the energy is dissipated through single or multiple Compton scattering or by the photoelectric absorption effect. The photoelectric absorption effect will produce on average, a distinguishable photo peak based on the energy of the gamma radiation that is absorbed. On the other hand, Compton scattering events will produce a broad distribution of number of scintillation light photons with no distinguishable photo peaks.
Within the context of positron emission tomography (PET), a radio labeled tracer is injected into a patient and preferentially retained by the cells of interest in order to emit positrons. Positrons emitted from the tracer normally travel over a distance of a few hundred microns in the tissues while losing kinetic energy. Each positron finally interacts with an electron of the medium resulting in an annihilation reaction where the masses of the two particles are transformed into two gamma photons or annihilation photons that have the exact same energy and travel in geometrically opposite directions.
Small animal PET scanners that are designed to image animals the size of rats and mice have been a driving force behind many of the advances of molecular imaging and have allowed characterization and understanding of some biological processes at the molecular level. The use of mice as animal models for applications in pharmacology, genetics, pathology and oncology, demand preclinical PET scanners that feature high spatial resolution and high sensitivity in order to visualize subtle distributions and to quantify low concentrations of PET tracers.
Advances in spatial resolution and sensitivity performance of imaging systems can open up applications currently out of the range of conventional PET scanners because of resolution limitations, such as mouse brain imaging and early lesion and metastasis detection in mouse models of cancer. Therefore, some of the most important research goals for preclinical PET imaging technology have been producing scanners with high sensitivity and high resolution.
Although PET can be a powerful imaging technique that has many applications in medicine, such as clinical oncology and pre-clinical pharmacology, the limited spatial resolution and sensitivity of PET scanners has suppressed its potential in small animal studies. Unfortunately, efforts to increase sensitivity and spatial resolution by using longer and narrower crystals for gamma radiation detection also eventually lead to degradation of the spatial resolution. Inter-crystal scatter (ICS) events are also a major source of error that leads to noise and degrades spatial resolution. A few methods have been developed to reduce the error associated with ICS events, but they typically require costly and demanding hardware and computational efforts that are not available for conventional Anger logic detectors.
The spatial resolution of conventional pixelated scintillation detectors is determined by the cross section of the scintillator crystal elements. The sensitivity can be increased by employing a compact system geometry to maximize the solid angle coverage, and by using long crystals for higher 511 keV gamma photon detection efficiency.
Unfortunately, long and narrow crystals in a small diameter gantry lead to increased penetration of oblique incident gamma rays before interaction. This causes event mispositioning, also called parallax error, degrading the spatial resolution uniformity and distorting the appearance of the reconstructed image of the source. Therefore, detectors with the capability of encoding the depth of annihilation photon interaction (DOI) are necessary. Much effort has been devoted to developing DOI PET detectors over the past several years. Among those designs, phoswich detector approaches obtain DOI information by measuring differences in light decay time between multiple layers of different scintillators. The phoswich detector design has attracted considerable interest and has been employed in several prototype scanners and commercial systems. Improved spatial resolution uniformity has been achieved in these phoswich DOI scanners compared to scanners of single layer design with equivalent scintillator volume and no DOI capability.
Inter-crystal scatter (ICS) events, where the incoming annihilation photons interact with more than one detection element within the same block detector, is another cause of event mispositioning in addition to parallax error. As the detection elements become narrower and longer, the fraction of these ICS events increases. With conventional PET detector designs that employ Anger logic positioning schemes, such ICS events appear as inaccurate detections. The spatial coordinates corresponding to the energy weighted mean of the multiple interaction sites are different from the location of first interaction. This error in determining the initial interaction location reduces image contrast and degrades spatial resolution. This leads to degradation of the lesion detectability and quantitative characteristics of an imaging system. Therefore, appropriate ICS event identification and correction methods are needed. Studies have shown that the capability of rejecting ICS events, or estimating the first interaction site of an ICS event using selection criteria, or maximum likelihood based on Compton kinematics, yields improved image quality and quantification. However, those approaches require complicated and costly data acquisition systems for measuring individual interactions of the ICS events and significant computational efforts for determining the location of first interaction, neither of which are available for conventional Anger logic detectors.
Accordingly, there is a need for improved scintillator detectors and imaging devices that have high spatial resolution and high sensitivity and are capable of high-resolution, high performance imaging.